Method and system for optical coherence tomography including obtaining at least two two-dimensional images of an object in three-dimensional space

ABSTRACT

A method and corresponding system for optical coherence tomography includes optical coherence tomography equipment that acquires at least two two-dimensional initial images of an object in planes of the object that are spaced apart from one another, in particular running parallel to one another, wherein the initial images each include a plurality of initial image values. In order to assure an examination of the object in the case of medical applications that is as reliable as possible, an interpolation of the initial image values of the at least two two-dimensional initial images is performed in three-dimensional space, wherein interpolation values are obtained that form a two-dimensional final image.

The present invention relates to a method and a corresponding system foroptical coherence tomography.

Optical coherence tomography (OCT) is a method of measuringlight-scattering specimens on their inside. Due to its light-scatteringproperties biological tissue is particularly suitable for diagnosticexamination by means of OCT. Since for OCT relatively low lightintensities are sufficient and the wavelengths of the light used mostlycome within the near infrared range (750 nm to 1350 nm), unlike ionisingX-ray diagnostics it does not contaminate biological tissue withradiation. It is therefore particularly significant for medicine and isroughly comparable to ultrasound diagnostics, wherein with OCT, light isused instead of sound. The running times of the light reflected ondifferent boundary layers within the specimen are recorded with the aidof an interferometer. With OCT, typically resolutions higher by one totwo orders of magnitude are to be achieved than with ultrasound, but themeasuring depth achievable is considerably smaller. Due to opticalscattering the cross-section images obtained usually only reach into thetissue up to a depth of a few millimeters. The currently most importantareas of application of OCT are in ophthalmology, dermatology and thediagnosis of cancer. However, there are also some non-medicalapplications, such as e.g. in materials testing.

In particular in medical applications of OCT special demands are placedon methods and systems to ensure an examination of the object which isas reliable as possible.

The object of the present invention is to specify a method as well as acorresponding system for optical coherence tomography that permits areliable examination of an object.

The aforesaid object is achieved by the method and the system accordingto the independent claims.

In the context of the inventive method for optical coherence tomography,at least two two-dimensional initial images of an object are acquired bymeans of an optical coherence tomography equipment in planes of theobject that are spaced apart from one another, in particular runningparallel to one another, wherein the initial images each comprise aplurality of initial image values, in particular intensity values.Furthermore an interpolation of the initial image values of the at leasttwo two-dimensional initial images in three-dimensional space isperformed, wherein interpolation values are obtained that form atwo-dimensional final image.

The inventive system for optical coherence tomography comprises anoptical coherence tomography equipment for the acquisition of at leasttwo two-dimensional initial images of an object in planes of the objectthat are spaced apart from one another, in particular running parallelto one another, wherein the initial images each comprise a plurality ofinitial image values, in particular intensity values, and said system ischaracterized by a processing device for the interpolation of theinitial image values of the at least two two-dimensional initial imagesin three-dimensional space, wherein interpolation values are obtainedthat form a two-dimensional final image.

The invention is based on the concept of initially acquiring two initialsection images in two parallel running planes of the object with theoptical coherence tomography equipment, and to then interpolate theinitial image values of the two initial section images inthree-dimensional space, which is spanned by the two initial sectionimages that are spaced apart from one another, in such a way that atwo-dimensional final section image is obtained from interpolationvalues, i.e. interpolated initial image values. For the derivation ofthe interpolation values initial image values of not only one but bothinitial section images are hereby interpolated. In so doing, aninterpolation value is derived from at least one initial image value ofa first initial section image, and at least one initial image value froma second initial section image. For example, an interpolation value isderived from eight initial image values, wherein four initial imagevalues originate from the first initial section image and four initialimage values originate from the second initial section image.

In the case of the OCT images obtained with the inventive method andsystem, cavities or other structures of a scale larger thanapproximately 10 μm appear more clearly than is the case in therespective initial section images. However, also smaller structures canbe identified and, if necessary, examined more reliably and faster.Particularly in the area of dermatology certain diagnostically relevantand interesting structures of the skin can as a result be recognized andexamined with particularly high reliability.

Preferably the respective initial section images, whose initial imagevalues are determined by inventive means, are acquired simultaneously oralmost simultaneously, in particular temporally spaced less than 40milliseconds apart, preferably less than 20 milliseconds.

Initial section images are simultaneously acquired, for example, duringan acquisition of initial images in the form of depth sections,so-called slices, in a first operating mode, which is described infurther detail below.

The almost simultaneous acquisition of initial section images, for thepurposes of the invention, takes place, for example, during anacquisition of initial images in the form of two-dimensional tomograms,so-called en-face images, in a second operating mode, which is describedin further detail below.

A simultaneous or nearly simultaneous acquisition of initial sectionimages can however also be performed by acquiring a completethree-dimensional data set of the object in a third operating mode,which is described in further detail below.

The simultaneous or almost simultaneous acquisition of the initialsection images to be averaged has the advantage that the effect ofpossible temporal changes of the object being examined on the finalimage, which is obtained by the averaging process, is eliminated orsignificantly reduced, which increases the accuracy and reliability ofthe images that are obtained during the examination of the object.

Preferably the initial image values of the at least two two-dimensionalinitial images are located on a regular grid in three-dimensional space,wherein adjacent initial image values in all three spatial dimensionsare equally spaced apart, in particular between approximately 2 μm and 4μm. As a result the inventive interpolation of the initial image valuesof the two initial images can be realized in a straightforward manner.

Further preferred is that the initial image values of the at least twotwo-dimensional initial images are interpolated by means of a trilinearinterpolation and/or a tricubic interpolation. The trilinear or tricubicinterpolation concerns a method for the interpolation within athree-dimensional regular grid, i.e. a grid with the same grid constantin all three spatial directions, wherein an interpolation value locatedin the center point of the respective grid cells of the grid isdetermined through linear or cubic interpolation from respectively eightinitial image values, which are located at the eight corners of the gridcell. The initial image values located at the eight corners of the gridcell represent the next neighbors of the respective interpolation value,which is why the trilinear or tricubic interpolation can also beregarded as a next neighbor interpolation. Final images of particularhigh diagnostic assessment value are obtained by means of this type ofinterpolation.

A quadratic interpolation can also be performed instead of or inaddition to a linear or cubic interpolation of the initial image valuesof next neighbors. In this case the initial image values of the at leasttwo two-dimensional initial images are interpolated using a triquadraticinterpolation.

Preferably the at least two two-dimensional initial images of the objectare real time images that are acquired at a rate of at least one imageper second, preferably at least five images per second. This isadvantageously possible, since the inventive interpolation represents astraightforward and very fast method that can be applied withouttemporal delay to OCT images acquired in real time. In this mannerinitial images acquired in real time can also be rendered in real time,i.e. with a corresponding image repetition rate, in the form ofinventively interpolated final images, with diagnostic information thatcan correspondingly be comprehended better, on a corresponding displaydevice.

In an additional advantageous embodiment of the invention the at leasttwo two-dimensional initial images of the object are acquired, in afirst operating mode, in planes of the object that are spaced apart fromone another, where light reflected or backscattered by the object isdetected only by a partial surface, in particular by two adjacent lines,of a spatially resolving detector of the optical coherence tomographyequipment, while the optical distance of a reflector from a beamsplitter of the optical coherence tomography equipment is changed by anoptical path that is significantly larger, in particular at least 100times, than the mean wavelength of light injected into the opticalcoherence tomography equipment. This operating mode permits theacquisition of initial images in the form of depth sections, so-calledslices, at high speed and consequently in real time, i.e. at a rate ofat least one image per second, in a straightforward manner and with highreliability. The increased diagnostic assessment value of the real timeimages captured and interpolated in this manner enables an even morereliable examination of the object.

In a further, likewise preferred, embodiment of the invention the atleast two two-dimensional initial images of the object are acquired, ina second operating mode, in planes of the object that are spaced apartfrom one another, where during a changing of the optical distance of areflector from a beam splitter of the optical coherence tomographyequipment the light reflected from the object is detected several times,in particular at most five times, by detector elements of a detector,wherein the change of the optical distance of the reflector from thebeam splitter is at most forty times the mean wavelength of lightinjected into the optical coherence tomography equipment. In thisoperating mode initial images in the form of two-dimensional tomograms,so-called en-face images, can be acquired with a high repetition rate,in particular in real time. The increased diagnostic assessment value ofthe real time images captured and interpolated in this mannerfacilitates, in the context of this embodiment also, an even morereliable examination of the object.

In this case the two planes of the object run preferably at differentdepths above and below a mean depth in the object. In particular the twoplanes of the object run at the same spacing above and below the meandepth in the object. Preferably the spacing of the two planes relativeto one another corresponds hereby to the same spacing of the initialimage values in all three spatial directions. These measures assure thedetermination of the interpolation values with particularly highreliability.

In a further advantageous improvement of the invention the mean depth orthe different depths of the planes running above and below the meandepth in the object is (are) set by the distance of the reflector fromthe beam splitter. Preferably the optical distance of the reflector fromthe beam splitter of the optical coherence tomography equipment ishereby changed by an optical path that is significantly larger, inparticular at least 100 times, than the mean wavelength of the lightinjected into the optical coherence tomography equipment. Theacquisition of the initial images in the form of en-face images atdifferent planes of the object can thereby be realized in astraightforward and quick manner.

In another preferred embodiment provision is made that prior toperforming the interpolation the plurality of original initial imagevalues of the at least two two-dimensional initial images is reduced atleast in one dimension, in particular in the direction of the depth ofthe object, by combining respectively at least two, preferably more thanten, original initial image values to one initial image value. Inparticular, the original initial image values concern sampling valuesthat are obtained by the successive sampling of an interference patternthat is obtained from different depths within a depth range of theobject. In this process preferably those original initial image valuesare combined that are obtained from the range, in particular the depthrange, of the object whose extent in the at least one dimensioncorresponds to the resolution, in particular the axial resolution ordepth resolution, of the optical coherence tomography equipment.

These measures on the one hand take into consideration that the samplingof the interference signal must be high enough so as not to violate theso-called sampling theorem. On the other hand the fact is thereby alsotaken into account that the spacing between two sampling values of theinterference signal is generally significantly smaller than the physicalresolution of the imaging optics of the optical coherence tomographyequipment. This means that several, preferably more than ten,sequentially following sampling values contain approximately the samephysical information, and can therefore be combined into one valuewithout significant loss of information. The initial image valuecorresponds thereby, for example, to a mean value or the median of theoriginal initial image values. As a result the captured data volume inthe form of the original initial image values is significantly reducedwithout significant loss of information, and furthermore the equalspacings of the initial image values in all three spatial directions,which are required for a particularly reliable execution of theinterpolation, can be straightforwardly implemented.

Furthermore preferred is that the depth of field of the imaging opticsof the optical coherence tomography equipment is larger than the spatialdistance, which preferably is identical in all three spatial directions,of the initial image values from one another. As a result the initialimages and the final images obtained through interpolation thereof arealways captured or obtained with the required accuracy.

Additional advantages, features and possible applications of the presentinvention are specified in the following description in the context ofthe figures. The drawings show:

FIG. 1 a schematic representation of an example of an optical coherencetomography equipment;

FIG. 2 a schematic representation of an example of a detector surfacefor illustrating a first operating mode;

FIG. 3 a spatial element of the object with cuts in first planes for theillustration of the first operating mode;

FIG. 4 a spatial element of the object with a cut in a second plane forthe illustration of a second operating mode;

FIG. 5 a spatial element of the object with cuts in second planes forthe illustration of the third operating mode;

FIG. 6 a) and b) two cross sections through the object and the samplearm of the interferometer for the illustration of the focus tracking;

FIG. 7 an example of a regular grid for the illustration of theinterpolation of initial image values;

FIG. 8 a schematic view for illustrating a sampling of an interferencepattern in the direction of the depth of an object in comparison to thephysical resolution in the direction of the depth;

FIG. 9 an additional schematic view for illustrating a compilation oforiginal initial image values, sampled in the direction of the depth ofan object, relative to respectively one initial image value incomparison to the physical resolution in the direction of the depth;

FIG. 10 an additional schematic view for the illustration of theinterpolation of the initial image values from two initial imagesobtained in the direction of the depth of the object;

FIG. 11 an additional schematic view for the illustration of theacquisition of the initial image values in one (left) or two (right)planes that are transversal to the direction of the depth of an object,as well as the interpolation of the initial image values of the initialimages obtained from the two planes (right).

FIG. 12 an example of an initial image (left) in comparison with acorresponding final image (right) that was obtained by means of thedescribed interpolation;

FIG. 13 a schematic representation of a system for implementing theinventive method for optical coherence tomography;

FIG. 14 a representation of a measuring head of the system;

FIG. 15 a monitor view for the illustration of the entry of patientdata;

FIG. 16 a monitor view for the illustration of the display of theentered patient data;

FIG. 17 a monitor view for the illustration of the adjustment of theskin moisture;

FIG. 18 an additional monitor view for the illustration of theadjustment of the skin moisture;

FIG. 19 a first monitor view for the illustration of the selection of asecond plane on the basis of a slice image;

FIG. 20 a second monitor view for the illustration of the selection ofthe second plane on the basis of the slice image;

FIG. 21 a third monitor view for the illustration of the selection ofthe second plane on the basis of the slice image;

FIG. 22 a fourth monitor view for the illustration of the selection ofthe second plane on the basis of the slice image;

FIG. 23 a fifth monitor view for the illustration of the selection ofthe second plane on the basis of the slice image;

FIG. 24 a sixth monitor view for the illustration of the display of aslice image stored in response to a user command, as well as theselection of the second plane on the basis of a temporarily stored sliceimage;

FIG. 25 a monitor view for the illustration of the selection of a firstplane for a slice image to be acquired on the basis of an en-face image;

FIG. 26 a monitor view for the illustration of the selection of slice aswell as en-face images that originate from a three-dimensional tomogram,in an image viewing mode;

FIG. 27 a monitor view for the illustration of an entry of comments inthe image viewing mode;

FIG. 28 a monitor view for the illustration of an administration mode ofthe system; and

FIG. 29 an example of an automatically generated examination report.

1. Optical Coherence Tomography Equipment

FIG. 1 shows a schematic representation of an example of an opticalcoherence tomography equipment, hereinafter also referred to as OCTequipment, with an interferometer 10, which comprises a beam splitter11, an illumination arm 12, a reference arm 13, a sample arm 14, and adetector arm 15. In addition, a radiation source 21 is provided forgenerating light, which is filtered by an optical filter 22 and isfocused through optics composed of lenses 23 and 24 onto an input region25 of an optical waveguide 26. The radiation source 21, together withthe optical filter 22, forms a device which is also designated as lightsource 20.

The light injected into the optical waveguide 26 is injected into theillumination arm 12 of the interferometer 10 by means of optics 28located in the output region 27 thereof. From there, the injected lightfirst reaches the beam splitter 11, through which it is forwarded intothe reference arm 13 and reflected by a movable reference mirror 16located at the end thereof and, after passing through the sample arm 14,illuminates an area 2 of a sample 1.

The light reflected, in particular backscattered, from the sample 1passes through the sample arm 14 once more, is superimposed in the beamsplitter 11 with the light from the reference arm 13 reflected at thereference mirror 16, and finally arrives via the detector arm 15 at adetector 30, which comprises a plurality of detector elements arrangedin a, preferably flat, surface and as a consequence, facilitates aspatially resolved detection of the light reflected from the sample 30or of a corresponding interference pattern due to the superpositionthereof with the light reflected at the reference mirror 16.

A CMOS camera is preferably used as the detector 30, the detectorelements (so-called pixels) of which are sensitive in the infraredspectral range, in particular in a spectral range between approximately1250 nm and 1350 nm. Preferably, the CMOS-camera has 512×640 detectorelements.

As the waveguide 26 a so-called multimode fibre is preferably used, thenumerical aperture and core diameter of which, for a specific wavelengthof the light injected into the fibre, allow not just one fibre mode tobe formed but many different fibre modes to be excited. Preferably, thediameter of the multimode fibre used is between approximately 1 mm and 3mm, and in particular approximately 1.5 mm.

The size of the illuminated area 2 on the sample 1 correspondsapproximately to the size of the illuminated area 17 on the referencemirror 16 and is defined firstly by the optics situated at the inputregion of the optical waveguide 26, which in the example shown comprisesthe lenses 23 and 24, and secondly by the optics 28 arranged in theoutput region of the optical waveguide 26.

In the described OCT equipment, the resulting interference pattern isdetected with the detector 30, wherein a corresponding interferencesignal is generated. The sampling rate of the detector 30 for samplingthe interference signal must be selected such that the temporalvariation of the interference pattern can be detected with sufficientaccuracy. In general this requires high sampling rates, if high speedsare to be achieved for a depth scan.

A depth scan is preferably realized in the system described by causingthe optical distance from the reference mirror 16 to the beam splitter11 to be changed with a speed v during the detection of the lightreflected from the sample 1 with the detector 30, by an optical pathlength which is substantially larger than the mean wavelength of thelight injected into the interferometer 10. Preferably, the lightreflected in at least 100 different depths of the sample 1 is therebycaptured by the detector 30. In particular, it is preferred that theoptical path is changed periodically with an amplitude which issubstantially larger than the mean wavelength of the light injected intothe interferometer 10. The change of the optical distance of thereference mirror 16 by the optical path or the amplitude respectively,is preferably at least 100 times, in particular at least 1000 times,greater than the mean wavelength of the light injected into theinterferometer 10. Because of the large path lengths in this distancevariation, this movement of the reference mirror 16 is also referred toas macroscopic movement.

Since the individual periods of an interference pattern in general needto be sampled at multiple time points respectively, the maximum possiblescanning speed in the direction of the depth of the sample 1 isdependent on the maximum possible sampling rate of the detector 30. Whenusing fast detector arrays with high spatial resolution, i.e. a largenumber of detector elements per unit length, the maximum sampling rateis typically in the range of approximately 1 kHz. For a mean wavelengthof the light injected into the interferometer of, for example, 1300 nm,this will result in a maximum speed for the depth scan of approximately0.1 mm/s, if four points per period of an interference structure aresampled.

To increase the speed of the depth scan, in the present OCT equipmentthe temporal profile of the sensitivity of the detector 30 for the lightto be detected is modulated with a frequency that is up to 40% greaterthan or less than the Doppler frequency f_(D), wherein the Dopplerfrequency f_(D) is related to the mean wavelength λ₀ of the lightinjected into the interferometer 10 and the speed v of the movingreference mirror 16 as follows: f_(D)=2v/λ₀. Typical frequencies of thismodulation are in the range between 1 kHz and 25 kHz. It is particularlypreferred that the frequency of the modulation of the detectorsensitivity is not equal to the Doppler frequency f_(D).

The light reflected by the sample 1 and impinging on the detector 30 issuperimposed with the modulated sensitivity of the detector 30, so thatduring the detection of the interference pattern impinging on thedetector 30, instead of a high-frequency interference signal with aplurality of periods, the detector 30 generates a low-frequency beatsignal which has markedly fewer periods than the high-frequencyinterference signal. In sampling this beating, considerably fewersampling time points per time unit are therefore necessary, withoutlosing any relevant information, than for sampling of the high-frequencyinterference signal without the modulation of the sensitivity of thedetector 30. For a given maximum sampling rate of the detector 30, thismeans that the maximum speed for a depth scan of the system can beincreased many times.

The sensitivity of the detector 30 can be modulated, e.g. directly orwith a controllable electronic shutter arranged in front of the detector30. As an alternative or in addition, properties of an optical elementin front of the detector 30, such as e.g. the transmittance of adetector lens for the light reflected from the sample 1, can bemodulated. Compared to systems with a constant detector sensitivity thisincreases the scanning speed by a factor of 4 or even 8.

The speed of the movement of the reference mirror 16 is in a fixedrelationship to the frequency of the modulation of the sensitivity ofthe detector 30 and is in particular chosen such that an integral numberof sampling time points, preferably four sampling time points, fit intoone period of the resulting beating signal.

The beating signals sampled in this way need to be further processedprior to being displayed, since these signals still contain theinterference information. The essential information to be displayed isthe amplitude and depth position of the respective interference, but notthe interference structure itself. In order to do this the beatingsignal must be demodulated, by determining the so-called envelope of thebeating signal e.g. by Fourier or Hilpert transformation.

Since the phase of the beating signal is in general unknown, and thiscan also differ for different beating signals from different depths, adigital demodulation algorithm is used, which is independent of thephase. For sampling the interference signal with four sampling timepoints per period, the so-called 90° phase shift algorithms arepreferably used. This allows a fast demodulation of the beating signal.

Preferably, one period of the modulation of the sensitivity of thedetector 30 comprises two sub-periods, wherein during a first sub-periodthe detector is sensitive and during a second sub-period the detector isinsensitive to the light to be detected. In general, the first and thesecond sub-period are equal in length. However, it can be advantageousto choose a different duration for the first and second sub-period. Thisis the case, for example, when the intensity of the light emitted by thelight source 20, or injected into the interferometer 10, and/or of thelight reflected from the sample 1, is relatively low. In these cases thefirst sub-period can be selected such that its duration is longer thanthe duration of the second sub-period. In this way, even at low lightintensities, in addition to a high depth scanning speed, a highsignal-to-noise ratio, and thus a high image quality, is ensured.

Alternatively to the sensitivity of the detector 30, the intensity ofthe light injected into the interferometer 10 can also be temporallymodulated, wherein the remarks on the modulation of the detectorsensitivity described above, apply accordingly with regard to thepreferred embodiments and the advantageous effects.

The radiation source 21 preferably includes a spiral-shaped wire, whichis surrounded by a transparent casing, preferably made of glass.Preferably, the radiation source 21 is implemented as a halogen lightbulb, in particular a tungsten halogen bulb, where a tungsten filamentis used as wire and the inside of the casing is filled with gas, whichcontains a halogen, e.g. iodine or bromine. By application of anelectrical voltage, the spiral wire is made to glow, which causes it toemit spatially incoherent light. The term spatially incoherent lightwithin the context of the present invention is to be understood as lightwhose spatial coherence length is less than 15 μm, and in particularonly a few μm, i.e. between approximately 1 μm and 5 μm.

The spatially incoherent light generated by the radiation source 21passes through the optical filter 22, which is implemented as aband-pass filter and essentially only transmits light within aspecifiable spectral bandwidth. The optical filter 22 has a bell-shapedor Gaussian spectral filter characteristic, wherein only those spectrallight components of the light generated by the radiation source 21 whichlie within the specified bandwidth about a mean wavelength of thebell-shaped or Gaussian spectral filter characteristic can pass throughthe optical filter 22.

A Gaussian spectral filter characteristic within the context of theinvention is to be understood to mean that the transmittance of theoptical filter 22 for light with particular wavelengths λ isproportional to exp[−[(λ−λ₀)/2·Δλ]²], where λ₀ designates the wavelengthat which the optical filter 22 has its maximum transmittance, and Δλ thestandard deviation, which is related to the full width at half maximum(FWHM) of the Gaussian transmittance curve as follows: FWHM≈2.35·Δλ.

A bell-shaped spectral filter characteristic is to be understood as aspectral plot of the transmittance of the optical filter, which can beapproximated by a Gaussian function and/or only deviates from a Gaussianfunction to the extent that its Fourier transform has essentially aGaussian shape with either no secondary maxima or only a small number ofvery low secondary maxima, the height of which is a maximum of 5% of themaximum of the Fourier transform.

The use of a radiation source 21 which a priori generates spatiallyincoherent light, in the detection of the light reflected by the sample1 by means of the two-dimensional spatially resolving detector 30,prevents the occurrence of so-called ghost images caused by coherentcrosstalk between light beams from different locations within the sample1 under test. The additional equipment for destroying the spatialcoherence, which is normally required when using spatially coherentradiation sources, can thereby be omitted.

In addition, thermal radiation sources such as e.g. incandescent orhalogen lamps can therefore be used to produce incoherent light, whichare much more powerful and more cost-effective than the frequently usedsuperluminescent diodes (SLDs).

Due to the optical filtering with a Gaussian or bell-shaped filtercharacteristic, the light generated by the radiation source 21 isconverted into temporally partially coherent light with a temporalcoherence length of preferably more than approximately 6 μm. This isparticularly advantageous with the described OCT equipment which is ofthe so-called time-domain OCT type, in which the length of a referencearm 13 in the interferometer 10 changes and the intensity of theresulting interference is continuously detected by means of a preferablytwo-dimensional detector 30 because, by filtering the light using thebandpass realized by the optical filter 22 on the one hand, a highlateral resolution of the image captured from the sample 1 is obtained,and on the other hand, due to the Gaussian or bell-shaped spectralfilter characteristic of the optical filter 22, the occurrence ofinterfering secondary maxima in the Fourier transform of theinterference pattern detected by the detector, which would cause theoccurrence of further ghost images, is avoided.

Overall, the described OCT equipment allows obtaining OCT images withhigh resolution and image quality in an easy way.

In the example shown, the optical filter 22 is arranged between theradiation source 21 and the optics formed from the two lenses 23 and 24on the input side. In principle, it is also possible however to providethe optical filter 22 between the two lenses 23 and 24 or between thelens 24 and the input region 25 of the optical waveguide 26.Essentially, an arrangement of the optical filter 22 is particularlyadvantageous if the light rays impinging on the optical filter 22 haveonly a small divergence, or in particular run parallel to one another,because, firstly, this reduces reflection losses at the boundarysurfaces of the optical filter 22 and secondly, it prevents any beamdisplacement due to light refraction. In the example shown therefore, anarrangement of the optical filter 22 between the two lenses 23 and 24 ofthe optics is preferred.

Alternatively or in addition, it is also possible however to mount theoptical filter 22 directly on the casing of the radiation source 21.This has the advantage that an additional filter component can bedispensed with.

Alternatively or in addition, it is also possible however to arrange theoptical filter 22 between the output region 27 of the optical waveguide26 and the illumination arm 12, for example in front of or between thelenses of the optics 28 located between the output region 27 of theoptical waveguide 26 and the input of the illumination arm 12.

In a simple and highly reliable variant the optical filter 22 comprisesan absorption filter, in particular a so-called dyed-in-the-mass glass,and an interference filter, wherein multiple, preferably between about30 and 70, thin layers with different refractive indices are applied tothe dyed-in-the-mass glass, for example, by vapour deposition, whichresults in an interference filter.

For the case where the optical filter 22 is integrated into the casingof the radiation source 21, the optical filter 22 is preferablyimplemented by applying such interference layers to the casing. As analternative, or in addition, it is also possible however to provide oneor more of the lenses 23, 24 or the lenses of the optics 28 with acorresponding interference filter.

2. Operating Modes of the OCT Equipment

The described OCT equipment can be operated in three different operatingmodes. The operating modes entail two real time modes, where OCT imagesof sample 1 are generated at a high rate of at least one image persecond, preferably approximately 5 to 10 images per second, as well asone static operating mode.

In the first operating mode, real time mode 1, two-dimensional depthsections of sample 1 are generated (so-called slices). This is realizedby using a CMOS camera as the detector 30, which permits the adjustmentof a so-called window of interest (WOI), where only a partial surface ofthe detector 30 is sensitive to light and converts the same tocorresponding detector signals. The reduction of the sensitive camerasurface is associated with a distinct increase in camera speed, so thatwith this setting more camera images can be generated per second than inthe full-image mode.

In the real time mode 1 a WOI is preferably selected that matches theentire camera length or width (for example 640 pixels) along onedirection, and has the—determined by the type of respective camera—leastpossible number of pixels (for example 4 pixels) in the other direction.As a result the speed of the camera is increased to such an extent thatOCT images can be acquired in real time.

This is preferably achieved with the previously described modulation ofthe sensitivity of the detector 30 or the modulation of the intensity ofthe light injected into the interferometer 10, or the light emitted bythe interferometer 10.

By way of example, FIG. 2 shows a detector 30 with a detector surfaceA1, which comprises a first plurality N1 of detector elements 31arranged in a plane, and has a length c1 and a width b1. With thesetting of a WOI as stated above, light is only detected by the detectorelements 31 that are located in a partial surface A2 of the detectorsurface A1, and converted into corresponding detector signals. Thesecond plurality N2 of the detector elements 31 of the partial surfaceA2 is smaller than the first plurality N1 of the detector elements 31 ofthe entire detector surface A1. The lengths c1 and c2 of the detectorsurface A1 or partial surface A2 are equal in size, while the widths b1and b2 of the detector surface A1 or partial surface A2 differ.

In the shown example the partial surface A2 is only 4 pixels wide, whilethe detector surface A1 is 512 pixels wide. The sensitive surface of thedetector surface A1 is consequently reduced by a factor of 128, whichsignificantly shortens the time duration required for the detection ofthe interference patterns and their conversion to corresponding detectorsignals.

As displayed in FIG. 3, only four (corresponding to the four pixel rowsof the partial surface A2) two-dimensional depth sections S (so-calledslices) are obtained in this example from the observed spatial element Rof sample 1, instead of a full three-dimensional tomogram. Due to theslices that are obtained in the first operating mode, this mode is alsoreferred to as the slice mode.

For purposes of further illustration the left part of FIG. 3 shows amodel of the human skin, where as an example a plane of atwo-dimensional depth section or slice acquired in operating mode 1,preferably in real time, is delineated.

In the second operating mode, the real time mode 2, two-dimensionaltomograms F are generated—as displayed in FIG. 4—at a certain depth T ofthe observed spatial element R of sample 1, wherein the depth T can bearbitrarily selected. Hereby the entire detector surface A1 of thedetector 30 is used for the detection of the light reflected by sample 1and the conversion thereof into corresponding detector signals, whereinhowever only maximally five camera images in each case are consideredfor the calculation of a tomogram F. To that end the reference mirror 16is moved periodically in the interferometer 10 at a certain distance tothe beam splitter 11 at an amplitude of about 1 μm about this distance,while up to five camera images are acquired, which are then computedinto an OCT image. In this manner tomograms F can be generated at a highrepetition rate, in particular in real time. In comparison to themacroscopic movement of the reference mirror 16 described above themovement of the reference mirror 16 in this case is microscopic.

The depth T at which the tomogram F is obtained can be arbitrarilyselected by means of the macroscopic movement of the reference mirror16—if applicable in combination with focus tracking, which is describedin more detail further below, of the light that is focused at a certaindepth T in the sample by means of the sample optics that are located inthe sample arm 14.

On the basis of the two-dimensional cuts F obtained in the secondoperating mode, through the sample 1 in planes that run substantiallyperpendicular to the direction of the light impinging on the sample 1,the second operating mode is also referred to as en-face mode.

For purposes of further illustration the left part of FIG. 4 shows amodel of the human skin, where as an example a plane of atwo-dimensional tomogram or en-face image acquired in operating mode 2,preferably in real time, is delineated.

In a third operating mode, a static mode, a complete three-dimensionaldata set is acquired with the aid of the macroscopic movement of thereference mirror 16 in combination with focus tracking.

At a mean wavelength of the light that is injected into theinterferometer 10 in the range of, for example, 1 μm the optical pathlength or amplitude of the macroscopic movement of the reference mirror16 is at least approximately 0.1 mm, preferably at least approximately 1mm.

In contrast to the conventional microscopic amplitude of the referencemirror movement, which is on the order of magnitude of the meanwavelength of the injected light, i.e. of up to typically 1 μm, in thedescribed OCT equipment a macroscopic movement of the reference mirror16 on the order of magnitude of 0.1 mm up to several millimeters takesplace.

During the macroscopic linear movement of the reference mirror 16, thelight reflected by sample 1 is forwarded to the two-dimensional detector30 via the interferometer 10 and detected by said detector successivelyat several time points for a certain time duration, which corresponds tothe integration time of the detector 30, in each case, and convertedinto corresponding detector signals.

In order for the light reflected from the reference mirror 16 and fromthe sample 1 to be able to interfere, the so-called coherence conditionhas to be satisfied, which states inter alia that the respectivelyreflected light waves must have a constant phase relation relative toone another in order to be able to interfere with one another. Due tothe use of light with a very short coherence length of typically 10 μmor less, the condition of a constant phase relation is only satisfied atcertain depths or depth ranges of the sample 1, which is also referredto as coherence gate.

Each position of the reference mirror 16 during the macroscopic movementcorresponds thereby to a certain depth within the sample 1, or to adepth range about this certain depth for which the coherence conditionis satisfied, so that the light reflected by the reference mirror 16 andby the sample 1 can interfere.

In the case of a periodic movement of the reference mirror 16, bothhalf-periods of the periodic movement of the reference mirror 16 caneach be used for the acquisition of detector signals.

In this manner successive two-dimensional cuts are acquired by thedetector 30 at different depths of the sample 1. This is illustrated inFIG. 5, where—representative for a plurality of two-dimensional cuts—afirst, second and third two-dimensional cut F1, F2 and F3 respectivelythrough a spatial element R is displayed. Such a two-dimensional cut“propagates” synchronously with the macroscopic movement of thereference mirror 16 in direction “a” through the observed spatialelement R of the sample 1, without the same having to be moved.

Every cut F1, F2 and F3 is located at a depth T1, T2 and T3 respectivelyof the sample 1, at which depth the coherence condition is satisfied ineach case, so that the light reflected by the reference mirror 16 and bythe sample 1 can interfere. The macroscopic movement of the referencemirror 16 in combination with the successive two-dimensional detectionof the light reflected by the sample 1 therefore has the effect of athree-dimensional depth scan.

The combination, as described above, of the macroscopic linear movementof the reference mirror 16 on the one hand with the detection of thelight reflected by the sample 1 with a two-dimensional detector 30 onthe other, facilitates a straightforwardly implemented and quickacquisition of a complete three-dimensional data set of the desiredspatial element R1 of the sample 1.

Due to the macroscopic movement of the reference mirror 16 athree-dimensional tomogram is hereby obtained instead of just atwo-dimensional image at a certain depth. In the process the sample 1has to be no longer moved relative to the second interferometer 20 forthe acquisition of a three-dimensional data record. This makes thedescribed OCT equipment compact, reliable and straightforwardlyhandleable, so that the same is particularly suitable for in vivo use.

The left part of FIG. 5 shows, for further illustration, a model of thehuman skin, where as an example a spatial element is delineated, ofwhich in operating mode 3 a three-dimensional tomogram is acquired.

3. Focus Tracking

The OCT equipment described above is designed such that during a fullstroke, meaning the path length or twice the amplitude, of the movementof the reference mirror 16 an interference signal of sufficiently highintensity and high sharpness is always obtained. As a result of thefocus tracking, which is described hereinafter, assurance is furthermoreprovided that the interference signal as well as the sharpness of thedetected interference pattern are maximized for all depths in the sample1.

To that end, during the detection of the light that is reflected fromthe sample 1, the focus, meaning the focal point of the imaging opticsof the interferometer 10 that is located in the sample arm 14, isadjusted in such a manner that the location of the focus in the sample 1and the location of that plane in the sample 1 for which in case of areflection of light the coherence condition is satisfied andinterference occurs, are essentially identical at all times during theacquisition of a tomogram of the spatial element R of the sample 1. Thisis illustrated in what follows on the basis of FIGS. 6a and 6 b.

FIG. 6a shows the case where the focus f of the—here only shownsimplified as a lens—sample objective 14 a in the sample arm 14 islocated at a depth of the sample 1 that does not coincide with thelocation of the coherence gate K. As a result the cut, through sample 1,that was obtained within the coherence gate K at a depth Ti is notimaged exactly in focus onto the detector 30 (see FIG. 1), so thatinformation losses would have to be accepted while detecting theinterference.

In FIG. 6b , on the other hand, the case is displayed where the focus fof the sample objective 14 a is set such that it is located within thecoherence gate K at a depth Ti. This tracking of the focus f of thesample objective 14 a, corresponding to the depth Ti of the coherencegate K in each case, is referred to as focus tracking. In this mannerthe interferometer 10 is focused during the depth scan on the respectivelocation of the coherence gate K at different depths Ti of the sample 1,so that images with a high sharpness are obtained from any depth ofsample 1.

The maximum optical scan depth Tm specifies to what depth beneath thesurface of the sample 1 the coherence condition for constructiveinterference is satisfied, and corresponding interference patterns areobtained.

The sample objective 14 a, which is displayed in FIGS. 6a and 6b in asimplified manner, preferably comprises several lenses that can bemoved, individually and/or in groups, in the direction toward the sample1 or away from the same. To that effect a piezo-electric actuator, forexample, is provided, in particular an ultrasound piezo motor, which iscoupled with the sample objective 14 a or the lenses, and moves the samealong one or several guideways, in particular guide rods or guidegrooves.

The movement of the sample objective 14 a or the lenses takes placepreferably synchronously with the macroscopic movement of the referencemirror 16 in the interferometer 10 (see FIG. 1). In this manner thefocus f of the sample objective 14 a tracks the coherence gate G, whilethe latter traverses successively different depths T1, T2 and T3 of thesample 1, from which two-dimensional cuts F1, F2 and F3 (see FIG. 5) areacquired respectively with the aid of the detector 30.

The synchronization of the macroscopic movement of the reference mirror16 and the focus tracking on the one hand, in combination with atwo-dimensional detector 30 on the other, assures a particularlystraightforward and quick acquisition of a plurality of in-focustwo-dimensional image sections at different depths of the sample 1, andthereby the acquisition of a full three-dimensional image data set ofhigh image quality.

Since the interferometer 10 and the optical imaging in the sample arm 14are continuously matched to one another, the interference signalsdetected by the detector 30 are at a maximum for any depth in the sample1, so that a very high signal to noise ratio results. Assurance isthereby furthermore provided that the lateral resolution is optimal forall depths in the sample 1, because the focus f of the imaging is alwayslocated within the coherence gate K. As a result OCT images with afaithful detail rendering and a high contrast are obtained.

Advantageously the speed v2 of the movement of the lens or the lenses ofthe sample objective 14 a in the direction of the sample 1 is smallerthan the speed v1 of the movement of the reference mirror 16. Preferablya ratio v1/v2 of the speeds of the reference mirror 16 and the lenses ishereby selected, which is approximately equal to 2·n−1, or is up toabout ±20%, preferably up to about ±10%, around this value. The locationof the focus f and coherence gate G is hereby matched to one anotherwith particularly high reliability.

As a result of the previously described selection of the ratio v1/v2 ofthe speeds of the reference mirror 12 and the lenses 42, assurance isprovided that the coherence gate K and focus f are superimposed on oneanother during the depth scan over the entire depth range beingobserved. In the example above of a sample with an index of refractionof n=1.4, the ratio v1/v2 of the speeds is in the range of approximately(2·1.4−1)±20%, meaning between approximately 1.44 and 2.16, and ispreferably approximately 2·1.4−1=1.8.

4. Trilinear Interpolation

The OCT images obtained with the OCT equipment and method describedabove can undergo an interpolation for the further improvement of theidentification of diagnostic information, for example in the area ofdermatology for the further improved identification of cavities orbulges in the skin that have a size of larger than approximately 10 μm.

A particularly advantageous interpolation method, in the context of OCTimages, particularly real time images, obtained with the OCT equipmentand method described above, is the so-called trilinear interpolation,where the initial image values of at least two two-dimensional initialimages, which were acquired in planes of the object running parallel toone another, are interpolated in three-dimensional space, so that atwo-dimensional final image is obtained. This is explained in detail inwhat follows.

The trilinear interpolation relates to a method for the multi-variateinterpolation in a three-dimensional regular grid, i.e. a grid with thesame grid constant in all three spatial directions. This is illustratedusing the grid shown in FIG. 7 as an example. On the basis of aninterpolation of the initial image values located on the eight cornersC000 to C111 of a cube, an interpolation value located in the centerpoint C of the cube is derived in each case.

The respective initial image values originate from initial imagesacquired in different planes of the object. The initial image values arelight intensity values at different locations in the correspondingtwo-dimensional initial images. The initial image values, i.e. the lightintensity values, with the coordinates C000, C001, C011 and C010originate, for example, from a first real time image acquired in theoperating mode 1 along a first depth section S (see FIG. 3), and theinitial image values, i.e. the light intensity values, with thecoordinates C100, C101, C111 and C110, originate from a second real timeimage acquired in the operating mode 1 along a second depth section S(see FIG. 3), spaced apart therefrom at a distance of the grid constant.The initial image values with the coordinates C000, C010, C110 and C100originate, in an alternative example, from a first real time imageacquired in the operating mode 2 in the form of a first two-dimensionaltomogram F (see FIG. 4), and the initial image values with thecoordinates C001, C011, C111 and C101 originate from a second real timeimage acquired in the operating mode 2 in the form of a secondtwo-dimensional tomogram F (see FIG. 4), spaced apart therefrom at adistance of the grid constant.

An identical resolution in all three spatial dimensions is selected fora trilinear interpolation of the OCT images, in particular the real timeimages, obtained with the OCT equipment and method described above.

This cannot be achieved without loss of resolution with OCT systemsknown from prior art because usually only a relatively high axial (i.e.longitudinal, in the direction of the light impinging on the object)resolution can be realized, while the lateral (i.e. the transverse,perpendicular to the direction of the light impinging on the object)resolution is usually considerably lower. A selection of equalresolution in all three spatial directions would therefore only bepossible by lowering the axial resolution, which as a rule however isnot desirable because of the large loss of information, since smallobjects can in that case no longer be resolved. In addition it is notpossible with OCT systems known from prior art to sample twotwo-dimensional images simultaneously, or at least almostsimultaneously. This applies particularly to en-face images and scanningsystems. A trilinear interpolation in real time is therefore almostimpossible because in that case movement artifacts also become relevant.

In contrast, in the case of the OCT images obtained with the OCTequipment and method described above, a trilinear interpolation ispossible in the case of the two-dimensional real time images captured inthe operating modes 1 and 2 (slice or en-face), as well as also for thepost-processing of the three-dimensional tomograms obtained in thestatic operating mode 3.

The axial (i.e. longitudinal) resolution is determined, in the case ofthe OCT equipment described above, primarily by the spectral bandwidthof the light source 20 and the index of refraction of the examinedobject 1, while the lateral (i.e. transverse) resolution is determinedprimarily by the optical imaging and the size of the detector elements31 of the detector 30 (see FIGS. 1 and 2).

The OCT equipment described above is tuned in such a way that lateraland axial resolution are almost equal and very high. Preferably theresolution in all three dimensions is approximately 3 μm×3 μm×3 μm.

For the lateral resolution this is achieved in particular through thefocus tracking described above, and for the axial resolution inparticular through the use of a light source 20, which comprises ahalogen lamp as a radiation source 21 in combination with a Gaussianfilter 22.

Furthermore preferred is that the depth of field of the imaging optics,in particular the sample objective 14, of the interferometer 10 (seeFIG. 1) is larger than the “grid spacing” of the initial image values,i.e. the spatial distance of the initial image values in the threedimensions. This provides assurance in every case that the initial imagevalues are always captured with high accuracy.

Preferably the fact is furthermore taken into account that the samplingof the interference signal must be high enough so as not to violate theso-called sampling theorem. This is explained in detail hereinafter.

FIG. 8 shows a scheme for the illustration of the sampling of aninterference pattern 40 in the direction of the depth T of an object, incomparison with the physical resolution 41 in the direction of depth T.With the OCT equipment and method described above, four points 42 eachare preferably sampled per interference period of the interferencepattern 40. An interference period in this case is the length of a half(mean) wavelength of the light injected into the interferometer (at amean wavelength of approximately 1.3 μm this corresponds toapproximately 0.65 μm). Consequently the distance 43 between twosampling points 42 is approximately 0.163 μm. The physical resolution 41in air is however approximately 4 μm. This means that approximately 24sequential lines in the depth direction T contain approximately the samephysical information, and can therefore be combined into one linewithout significant loss of information. This in turn has the effectthat the resulting volume image point (so-called voxel) has almost thesame extent in all three dimensions, meaning it correspondssubstantially to a cube. The initial image value corresponds thereby,for example, to a mean value or the median of the original initial imagevalues.

FIG. 9 illustrates the previously described combining of originalinitial image values, which were sampled in the direction of the depth Tof the object in several sequential lines 44, to one line of only oneinitial image value and one line height, i.e. a longitudinal extent 45in the depth direction T that corresponds to the lateral extent 46 of animage point (pixel) of the line, perpendicular to the depth direction T.

In operating mode 1, where slices are acquired in real time, twoneighboring lines of the detector 30 are simultaneously read out in thecase of the trilinear interpolation. In the example of the detector 30shown in FIG. 2 this means that the width b2 of the partial surface A2of the detector 30 is selected such that said width extends only acrosstwo detector elements 31 in the direction of the width of the detector30. The partial surface A2 in that case comprises only 2×640 detectorelements 31 that are successively read out during a macroscopic movementof the reference mirror 16 (see FIG. 1), and are computed into atwo-dimensional final image in the manner described above.

This is illustrated on the basis of FIG. 10. Two initial images S in theform of two depth sections (compare FIG. 3), which were acquired in thedirection of the depth T of the object, are combined to a final image S′using trilinear interpolation.

Since the two initial images S in the form of two depth sections areacquired simultaneously and within a very short time, it is assured thata possible relative movement between sensor head and object, inparticular the human or animal skin, is of no significance during theacquisition of the two two-dimensional initial images S.

In the operating mode 2, in which en-face images are acquired in realtime, the reference mirror 16 (see FIG. 1), which is located in a meanposition, performs only a microscopic, preferably oscillating, movementfrom approximately +/−5 μm to +/−40 μm. In this case the position oroptical imaging property of the sample objective 14 is preferably set insuch a way that a focal point thereof has a mean depth position that ispredefined by the macroscopic displacement of the reference mirror 16.In the case of the trilinear interpolation of the en-face imagesacquired in real time—in contrast to operation without trilinearinterpolation—two initial images in the form of two en-face images arecaptured each at two different positions of the reference mirror 16, andcomputed into a two-dimensional final image in the form of an en-faceimage.

This is illustrated on the basis of the diagram shown in FIG. 11, whichshows the progression of the position P of the reference mirror 16 overtime t.

In the left part of the diagram of FIG. 11 the case without trilinearinterpolation is displayed. In the operating mode 2 a two-dimensionalinitial image F is hereby obtained in the form of a tomogram from acertain depth in the object, by measuring at five positions P of thereference mirror 16, which positions are located symmetrically about amean position P₀.

In the right part of the diagram of FIG. 11 the application of thetrilinear interpolation is illustrated. Two two-dimensional initialimages F are obtained by measuring at five positions P of the referencemirror 16 in each case. The five positions P are located eachsymmetrically about the positions P₁ and P₂, which preferably arethemselves located symmetrically about the mean position P₀ of thereference mirror 16. The distance 47 between the positions P₁ and P₂ ofthe reference mirror 16 is in this case determined by the axial and/orlateral pixel size 45 and 46 respectively (see FIG. 9). With thepreferred symmetric location of the positions P₁ and P₂ thecorresponding tomograms F each are located in the object above and belowthe mean depth location by half a pixel size. The initial images Facquired in this manner then undergo a trilinear interpolation, duringwhich the final image F′ is obtained.

Preferably the depth of field of the optical imaging in theinterferometer 10 (see FIG. 1) is selected such that the same is largerthan half the voxel size. With a preferred voxel size of 3 μm, the depthof field must therefore be larger than 1.5 μm.

Since the described acquisition of the two initial images in theoperating mode 2 takes place immediately one after the other, typicallytemporally separated by about 0.014 seconds, the effect on the obtainedfinal image of a possible relative movement between sensor head andobject, in particular the skin, between the acquisition of the twooriginal initial images is almost completely ruled out or negligiblysmall.

During the acquisition of the images the sensor head is preferably indirect contact with the surface of the object to be examined, inparticular the skin, which significantly reduces the probability of arelative movement. This is a particular advantage with acquisitions ofimages of the human or animal skin, since the same is generally elasticand adheres to the tip of the sensor head, particularly during theapplication of a gel, so that slight lateral movements or the slighttipping of the sensor head most often do not lead to a relative movementbetween skin and sensor head.

In the operating mode 3, in which static three-dimensional tomograms areacquired, a beat—as described above in detail—is generated between thedetector sensitivity modulation on the one hand, and the interferencesignal to be captured on the other. As a result the distance between theindividual sampling points in the depth direction is larger than in theoperating mode 1, so that correspondingly fewer sample points,preferably between 6 and 10, in particular 8, are combined to maintain acube-shaped three-dimensional image element (voxel).

FIG. 12 shows an example of an initial image (left) in comparison with acorresponding final image (right) that was obtained by means of thedescribed interpolation. The final image is less noisy in contrast tothe initial image, and appears therefore “softer” or “smoother”. Duringcomparisons in the interpretation of the images for diagnosis purposesit has turned out, in particular in the field of dermatology, that therelevant diagnostic information in each case can be obtained faster andmore reliably from the final images obtained through trilinearinterpolation. This applies in particular to cavities or inhomogeneitieswith a size of typically more than 10 μm.

The explanations provided above regarding the trilinear interpolationapply correspondingly also to a tricubic interpolation, where theinitial values are not interpolated using a linear function, but insteadusing a cubic function.

5. System for Optical Coherence Tomography

FIG. 13 shows a schematic representation of a system 50 for implementingthe inventive method for optical coherence tomography. The system 50comprises a housing 51, entry devices in the form of a key pad 53, acomputer mouse 54 as well as a foot switch device 55 that has a left,center and right foot switch 55 l, 55 m and 55 r respectively. Thehousing 51 in the displayed example is designed to be mobile by beingprovided with rollers 56.

Furthermore a measuring head 57 is provided, which is connected to thehousing 51 via a cable 58 or a cable hose or -pipe. The measuring head57, in its idle position, is plugged into a measuring head holderprovided on or in the housing 51, from which said measuring head can beremoved during the acquisition of the OCT images, which in the figure isindicated by the measuring head 57, represented by a dashed line, andthe cable 58, represented by a dashed line.

The system has a display device 52 in the form of a flat panel monitorthat can display OCT images 60 and 61, which were captured by placingthe measuring head 57 on an object, in particular the skin of a patient.In the example shown in the figure the first OCT image 60 concerns adepth section running substantially perpendicular to the surface of theobject being examined, which depth section was acquired in the operatingmode 1 described above, and the second OCT image 61 concerns atwo-dimensional tomogram that runs substantially parallel to the surfaceof the object being examined, and that was acquired in the operatingmode 2 described above.

In the area of the first OCT image 60 a straight line 62 is displayed onthe display device 52, which straight line can be moved upward ordownward in the direction of the indicated double arrow, for example byselecting a corresponding position of the straight line 62 relative tothe first OCT image 60 with the aid of the entry devices 53, 54 and 55.The system 50 is configured in such a way that, corresponding to theselected location of the straight line 62 in the displayed first OCTimage 60, a plane in the object being examined, running perpendicular tothe displayed first OCT image 60, is determined automatically and atwo-dimensional tomogram is acquired there, which is then displayed asthe second OCT image 61.

The first OCT image 60 is preferably a so-called slice, while the secondOCT image 61 preferably represents a so-called en-face image, which hasbeen acquired in a plane corresponding to the straight line 62 in thefirst OCT image 60.

A depth selection display 63 in the form of a switch symbol, which ismovable along a straight line, is furthermore displayed on the monitorof the display device 52, which switch symbol shows the depth that wasselected through a selection of the location of the straight line 62relative to the first displayed OCT image 60. Alternatively or inaddition the depth can also be indicated in the form of numericalvalues.

One or several additional selection displays can furthermore be providedon the display device 52. In the displayed example a selection screen 64is provided that shows one or several properties of the object to beexamined. These properties are preferably selected and entered by anoperator prior to the acquisition of corresponding OCT images. In thecase of dermatological applications this concerns, for example, aparameter for the characterization of the moisture of the skin of therespective patient. In the corresponding selection screen 64 acorresponding switch symbol can then be moved continuously or inspecified steps along a straight line between the positions “dry skin”on the left and “moist skin” on the right.

The interferometer 10 displayed in FIG. 1, including the optics 28 andthe detector 30, is integrated in the measuring head 57. The lightsource 20, including the optics in the form of the two lenses 23 and 24on the input side, is preferably integrated into the housing 51 of thesystem 50. The optical waveguide 26, which couples the light source 20on the one hand and the interferometer 10 on the other with one another,is guided within the cable 58 from the housing 51 to the measuring head57 in this case. In cable 58 electrical lines are furthermore guidedthat on the one hand serve the purpose of supplying power andcontrolling the measuring head 57, and on the other conduct the detectorsignals, which are generated during the capture of OCT images, of thedetector 30 from the same into the housing 51, where they are suppliedto a processing device (not displayed).

The measuring head 57, which in FIG. 13 is shown only heavilyschematized, is displayed in detail in FIG. 14. A grip 57 b is providedin the lower area of a measuring head housing 57 a of the measuring head57, which grip can be used by an operator to remove the measuring head57 from the measuring head holder on or in the housing 51, or to plugsaid measuring head again into the measuring head holder, and to placesaid measuring head onto the object during the acquisition of OCT imagesand, if applicable, to guide said measuring head along said object. Inthis context the measuring head 57, with a contact surface 57 c that islocated on the front end of the measuring head housing 57 a, is broughtinto contact with the object to be examined, in particular the skin of apatient.

In the center of the contact surface 57 c a window 57 d is provided,through which light from the sample arm 14 of the interferometer 10 (seeFIG. 1) located in the measuring head 57 can pass, and can therebyirradiate the object to be examined. The light reflected and/orbackscattered at different depths of the object reenters the sample arm14 of the interferometer 10 through the window 57 d and can there becaptured and analyzed in the form of interference phenomena, as wasalready illustrated in detail above.

A status display device 57 e, preferably in the form of an indicatorlight, is furthermore provided on the measuring head housing 57 a, bymeans of which for example the readiness of the system 50 and/or themeasuring head 57 for the capturing of OCT images is shown.

The cable 58, which can also be designed as a cable conduit or hose, isconnected to the measuring head 57 in the area of the rear end of themeasuring head housing 57 a.

With the system 50 for optical coherence tomography described above,three- and two-dimensional cross section images of an object, inparticular the human skin, can be acquired, where penetration depthsinto the human skin of up to approximately 1 mm can be reached, and thesize of the surface of the skin area examined has typical dimensions ofapproximately 1.8×1.5 mm. Due to the infrared radiation that is used inthe described system 50, with a preferred mean wavelength ofapproximately 1.3 μm, radiation exposure of the patient, such as forexample during the use of x-rays, can be ruled out. The OCT imagescaptured with the described system 50 furthermore have high resolutionand permit a display of individual object structures on a scale down to3 μm. Not least, the OCT images captured with the system 50 can be usedfor measuring the absolute geometric extent of the different structures,i.e. their size.

The system 50 has—even if not explicitly shown—a control device for theinventive control of the system 50, in particular the optical coherencetomography equipment, or the execution of the sequences describedpreviously and hereinafter. The system furthermore comprises aprocessing device for the processing of different data, including theinterpolation of initial image values described above. The controldevice and/or the processing device are preferably integrated intohousing 51 of the system 50.

6. Workflows, Depth and Lateral Navigation

The functionality and the handling of the system 50 for opticalcoherence tomography are described hereinafter, by way of example, onthe basis of typical and/or preferred sequences (so-called workflows).The advantages hereby achieved are also explained.

FIG. 15 shows the content of the monitor 70 of the display device in theadministration mode, in which the system is automatically set afterstarting. A status display 71 in the form of a suitable symbol, forexample a green disc, shows the readiness of the system. Preferably thereadiness of the system, in particular for the acquisition of OCTimages, is shown simultaneously through the activation of the statusdisplay 57 e provided on measuring head 57. Hence, an operator has theopportunity of recognizing the readiness of the system solely on thebasis of the status display 71 on the monitor 70, or on the basis of thestatus display 57 e on the measuring head 57.

Information regarding the object to be examined, in particular regardinga patient, can be entered in an entry field 72. Preferably the system ishereby configured in such a way that a acquisition of OCT images is onlypossible when at least one item of the information requested in theentry field 72 has been entered, for example at least the last name of apatient.

The information entered in entry field 72, in particular the first andlast name, patient identification number as well the data of birth, thenappear—as illustrated as an example in FIG. 16—in the correspondingfields 72′ in the upper area of the monitor display 70.

As soon as the measuring head 57 is removed from the measuring headholder that is located on or in the housing 51 of the system, the systemstarts automatically in the operating mode 1 (so-called slice mode). Anoptical gel is applied to the contact surface 57 c of the measuring head57 prior to bringing the contact surface 57 c of the measuring head 57in contact with the skin of the patient, which gel assures that, on theone hand, sharp index of refraction transitions between the skin and thewindow 57 d of the measuring head 57 are bridged (so-called indexmatching) and, on the other hand, irregularities on the skin surface areevened out. Preferably the amount of the optical get applied to thecontact surface 57 c is, depending on the application case, betweenapproximately 2 μl and 10 μl.

After application of the gel, the contact surface 57 c of the measuringhead 57 is pressed against the patient's skin area to be examined andmoved back and forth slightly over the skin area by an operator in orderto achieve an optimal distribution of the optical gel.

Since the system is already in the slice mode immediately after theremoval of the measuring head 57 from the measuring head holder, a sliceimage 73 is captured immediately after a contact with the skin area tobe examined is established, and displayed in the area of the center ofthe monitor 70, as illustrated in FIG. 17. A display 74 is provided inthe right area of the monitor 70, where the skin type of the skin beingexamined in each case is set or shown. Preferably this relates to aparameter that characterizes the moisture content of the skin area beingexamined. A corresponding switch symbol can in this case be moved by theoperator, in steps or also continuously, on a scale between “dry skin”and moist skin”. The selection of this parameter specifies the ratio ofthe speeds with which the reference mirror 16 and the lens or lenses ofthe sample objective 14 a (see FIG. 1 as well as FIGS. 6a and 6b ) move,in order to assure an optimal focus tracking.

For the acquisition of the slice image 73 that is displayed in FIG. 17 aposition of the switch symbol of the display 74 was selected that islocated slightly above the center of the scale, corresponding to a moremoist skin. The result is a bright and relatively high-contrast sliceimage 73.

In contrast, the slice image 75 displayed in FIG. 18 was captured with aparameter setting where the switch symbol of the display 74 is locatedbeneath the center of the scale, which corresponds to a more dry skin.The contrast of the slice image 75 captured with this setting issignificantly lower compared with the slice image 73 shown in FIG. 17,as can be clearly seen in FIG. 18. The reason is that the focus of thesample objective 14 was not or not always positioned in the range of thecorresponding coherence gate during the acquisition of the slice image75 at the different depths of the skin. For further details reference ismade to the explanations above in the context of focus tracking.

Proceeding from a slice image 73 (see FIG. 17) obtained with the optimalsetting of the parameter corresponding to the skin moisture, actuationof a corresponding switch, preferably by means of a prolonged pressingof the center foot switch 55 m (see FIG. 13), enables switching from theslice mode to the en-face mode, in which—as illustrated in FIG. 19—theslice image 73 is displayed scaled-down (so-called thumbnail) in theright area of the monitor 70, and simultaneously an en-face image 76that was acquired in the operating mode 2, the so-called en-face mode,is shown in the center area of the monitor 70. The displayed en-faceimage 76 is preferably a real time image that is being acquired andupdated at a repetition rate of at least one image per second. Thescaled-down display of the slice image 73 in the right area of themonitor 70, on the other hand, is a static image, corresponding, forexample, to the last slice image acquired and displayed in the slicemode (see FIG. 17) in real time.

The depth of the skin at which the displayed en-face image 76 isacquired can be selected by the operator via a depth selection switch 77shown on the monitor 70, by actuating a corresponding switch symbol, forexample using the computer mouse 54, the key pad 53 and/or the footswitch device 55 (see FIG. 13). Preferably the setting or selection of aspecific depth takes place via the left foot switch 55 l, which isdesigned as a rocker switch and which initiates a change in depth towardgreater or smaller depth via actuation of the rocker toward the front orthe back.

The selection of a certain depth on the basis of a displayed first OCTimage, as described above, at which depth a second OCT image isacquired, is also referred to as depth navigation in the context of theillustration of the inventive system and method.

Preferably the system is configured such that a selection of the depthat which an en-face image is to be captured can be tracked with anaccuracy as low as one micrometer. Basically it is possible that thesize of the steps at which depth navigation is performed is specified.As an example, the specification can be made prior to the start of anexamination, i.e. the acquisition of several OCT images of a patient,that the selection of the respective depth for en-face images is to takeplace in steps of 5 μm. In this manner the depth navigation can beadapted individually to the respective purpose of diagnosis.

The previously described selection of a certain depth for theacquisition of an en-face image is explained hereinafter in more detailon the basis of the monitor displays shown in FIGS. 20 to 23.

FIG. 20 shows an en-face image 80 that was acquired at a depth betweenthe window 57 d positioned on the measuring head 57, which window can berecognized in the form of a horizontal line 79 in the correspondingslice image 81, and the skin surface, and for that reason shows only across section of the gel layer that is located between the window 57 dand the skin. The depth set in this example is shown by means of ahorizontal straight line 78 marked in slice image 81, which is displayedscaled-down. Furthermore the selected or set depth can also be obtainedfrom the respective position of the depth selection switch 77 and/or acorresponding numerical value display.

FIG. 21 shows an en-face image 32 that was acquired in a plane that islocated in the upper most region of the skin surface, as can berecognized from the position of the straight line 78 that serves as adepth display, relative to the slice image 81 that is displayedscaled-down. The region between the straight line 78 on the one hand andthe horizontal line 79 on the other, which is due to light reflectionson the window 57 d of the measuring head 57, corresponds to the gellayer located between the window 57 d and the skin.

By actuating the left foot switch 55 l or the corresponding switchsymbol in the depth selection display 77, for example with the aid ofthe computer mouse 54, the straight line 78 can be moved (see doublearrow) relative to the slice image 81, which is displayed scaled-down,with the result that planes located at different depths in the skin canbe selected, in which corresponding en-face images are acquired andshown on the monitor view.

The principle of depth navigation is further illustrated in the rightlower part of FIG. 21 on the basis of a plane that is marked in a skinmodel, said plane running substantially parallel to the skin surface andcan be moved to different depths in the direction of the double arrow.

FIG. 22 shows, as an example, an additional en-face image 83 that wasacquired at another depth of the examined skin area. The plane of theobtained en-face image 83 is now completely within the examined skinarea, as can be recognized from the position of the straight line 78relative to the slice image 81. Moreover, the statements in connectionwith the FIGS. 20 and 21 apply correspondingly.

FIG. 23 shows an advantageous use of the depth navigation describedabove for the detection of diagnostically relevant information. As anexample, the depth for the acquisition of an en-face image 84 can beselected in the displayed slice image 81 via the selection of thelocation of the straight line 78, in order to, for example, furtheranalyze a cavity 85, which is suspected on the basis of the slice image81, in a corresponding plane of the en-face image 84, which plane runsperpendicular to the slice image 81.

Additional aspects of the described sequence during the acquisition ofOCT image with the inventive system are explained hereinafter in moredetail.

In the right area of the monitor 70, which is shown in FIG. 24, ascaled-down slice image 85 is displayed that was acquired in the slicemode and stored in a non-volatile memory of the system, for example ahard disk memory, via actuation of a corresponding switch, preferablyvia briefly pressing the center foot switch 55 m (see FIG. 13).

A depth navigation, as it was explained in detail in the context of theFIGS. 19 to 23, can be performed on the basis of the stored anddisplayed slice image 85. Preferably the system is hereby configured insuch a manner that an additional slice image 86 is automaticallygenerated and displayed in the right area of the monitor 70, if theslice image stored last, in this case the slice image 85, is alreadyolder than a specified time duration, for example 10 seconds, during theswitch from the slice mode to the en-face mode. This case is shown inthe example displayed in FIG. 24, where the slice image 85 was acquiredat a first time point and stored after entry of a corresponding operatorcommand, and the switch to the en-face mode was only made after a timeduration of more than 10 seconds after the acquisition and storing ofthe slice image 85. In this case an additional slice image 86 wasacquired immediately after the switch to the en-face mode, storedtemporarily, for example in a volatile memory of the system, anddisplayed scaled-down in the right area of the monitor 70, where astraight line 78 is superimposed in the region of the OCT image 86 forperforming the depth navigation described above, on the basis of whichstraight line an operator can recognize and control at what depth in theobject, in each case, a corresponding en-face image 87 is acquired anddisplayed, preferably in the center area of the monitor 70.

This configuration of the system assures that the depth navigationpreviously described is always performed on a most current slice image,so that possible relative movements between the measuring head on theone hand and the object on the other, including movements within theobject itself, can be taken into account, and can therefore not affectthe reliability of the acquisition of OCT images, in particular en-faceimages, negatively.

The adjustable time interval between the acquisition and storing of aslice image on the one hand, and the change from slice mode to theen-face mode on the other, where due to the time interval being exceededan additional slice image is acquired, temporarily stored and shown onthe monitor 70, was set to 10 seconds in the displayed example.Basically it is however also possible to select this time interval to besignificantly shorter, for example 5 seconds, if the type of respectiveexamination requires this. This can be the case, for example, if themeasuring head, due to larger movements of the object, in particular apatient, can not be maintained sufficiently long in a fixed positionrelative to the object. On the other hand it is however also possible tospecify longer time intervals, for example 15 seconds, when, forexample, the object to be examined remains stationary for a longer timeinterval and a fixed relative position between measuring head and objectcan be assured.

FIG. 25 shows a monitor view 70 after the switch from the en-face mode,whose monitor display is displayed as an example in FIG. 24, back to theslice mode. In the right area of the monitor view 70 the slice image 85is shown, which in this case is permanently stored due to an operatorcommand, not, however, the slice image 86 (see FIG. 24), which is onlytemporarily stored for navigation purposes. Furthermore the en-faceimage 87 that was acquired and shown last in the en-face mode isdisplayed in a scaled-down form after the switch to the slice mode.

A currently acquired slice image 88 is displayed in the center area ofthe monitor 70. Analogously to the depth navigation described above, thesystem is configured in such a way that a plane, which is perpendicularto the plane of the displayed en-face image 87, can also be selected inthe en-face image 87, which is displayed scaled-down, with the aid of anadditionally displayed straight line 89, in which plane the slice image88 is acquired.

The selection, which is performed on the basis of an en-face image, ofslice image planes, which run substantially parallel to the lightimpinging on the object and perpendicular to the skin surface or to theplane of the en-face image, can also be referred to as lateralnavigation. Moreover, the statements above in connection with the depthnavigation apply correspondingly.

The principle of lateral navigation is further illustrated in the rightlower part of FIG. 25 on the basis of a plane that is marked in a skinmodel, which plane runs substantially perpendicular to the skin surfaceand can be moved laterally in the direction of the double arrow.

In the monitor view 70 shown in FIG. 25 a slice image 88′ is furthermoredisplayed in the right area, which was stored in the currently selectedslice mode via a corresponding operator selection command. Furthermore a3D symbol 90 is shown that indicates that in the meantime athree-dimensional tomogram acquired in operating mode 3 was alsoacquired and stored.

7. Image Viewing and Administration Mode

After the completion of the acquisition of one or a plurality of,possibly different, OCT images, the measuring head 57 is plugged againinto the measuring head holder located on the housing 51 of the system50, whereupon the monitor display 70—as shown in FIG. 26—transitionsautomatically into an image viewing mode, in which the operator canselect the displayed and stored OCT images 85, 87, 88′ and 90, which aredisplayed scaled-down in the right area of monitor 70, wherein eachselected scaled-down image is displayed enlarged in the center area ofthe monitor 70.

In the case where a three-dimensional tomogram 90 is selected, aperspective rendition of the acquired three-dimensional tomogram can beprovided in the center area of the monitor 70. For certain diagnosticapplications it can be advantageous, however, to display, in each case,a slice image 91, which originates from the three-dimensional tomogram,and an en-face image 92 together and enlarged in the center area of themonitor 70, as shown as an example in FIG. 26. It is hereby advantageousto use the principle of depth and lateral navigation, as describedabove, in this case also, wherein corresponding straight lines 93 and 94are superimposed on the displayed slice and en-face images 91 and 92respectively. The user can specify the plane of each displayed sliceimage 91 via the selection of a straight line 93 or 94 as well as viathe selection of the location of the selected straight line 93 or 94 inthe area of the en-face image 92. Furthermore a selection of a plane ofthe en-face image that is to be shown and that originates from thethree-dimensional tomogram can take place via selecting and sliding ofthe straight line 93 in the area of the slice image 91.

The capability of entering comments in the image viewing mode isillustrated on the basis of the monitor view 70 displayed in FIG. 27. Tothis end the operator initially selects an OCT image for commenting,which in the displayed example is the slice image 85, and then opens acorresponding comment field 97 that is associated with this image, inwhich comment field then any comments can be entered in the form of freetext. Furthermore a general comment field 96 is opened, in which acomment regarding the performed examination can be entered which isassociated with the entirety of the OCT images 85, 87, 88′ and 90acquired during this examination and which is shown during retrievals ofat least one of these images together with the retrieved image. Theslice image 85, which is selected and displayed scaled-down in the rightarea of the monitor 70, is displayed enlarged in the center of themonitor 70 shown in FIG. 27.

After completion of the analysis and, if applicable, the commentingregarding to the OCT images acquired during the examination, anadministration mode of the system 50 can be selected in which theperformed examination is shown in monitor view 70, as displayed in FIG.28, in the form of, respectively, one line 98. By selecting thecorresponding line 98, the operator can again switch into the imageviewing mode and analyze and, if applicable, enter comments for theacquired OCT images.

An examination report of the performed examination, as shown as anexample in FIG. 29, can be generated—automatically or after completionof the examination or after a user command. In the examination report,which is preferably created in the HTML format, the patient informationthat was entered prior to the examination, the OCT images 85, 87, 88′and 90 that were acquired during the examination and stored in responseto a user command, as well as the respectively entered comments 96 and97 are compiled in the form of an overview.

8. Additional Inventive Aspects of the System and Method

The OCT system and method previously described in more detail hasindividual features or feature combinations that make the system andmethod more straightforward, quicker and more reliable with regard tohandling and image acquisition, without all of the features listed inthe preamble and/or characterizing portion of the independent claimsbeing hereby imperatively required. These features or featurecombinations are likewise considered an invention.

In particular, a system for optical coherence tomography is consideredan invention with at least one interferometer for the emission of lightfor the irradiation of an object, and a detector for the detection oflight that is reflected and/or backscattered from the object, whereinthe system is characterized by one or a plurality of features, whichwere previously described in more detail, in particular in the sections1 to 7 and/or in connection with the FIGS. 1 to 29. The methodcorresponding to this system is likewise considered an invention.

Furthermore, a method for optical coherence tomography is considered aninvention, where a first image is acquired, in particular in real time,in the region of a first plane of an object by means of an opticalcoherence tomography equipment, and the first image is displayed on adisplay device, in particular as a real time image, wherein the methodis characterized by one or a plurality of features, which werepreviously described in more detail, in particular in the sections 1 to7 and/or in connection with the FIGS. 1 to 29. The system correspondingto this method is likewise considered an invention.

The invention claimed is:
 1. A method for performing optical coherencetomography, the method comprising the steps of: successive sampling ofan interference pattern at different depths within a depth range of anobject to obtain a plurality of original initial image values at thedifferent depths within the depth range of the object using opticalcoherence tomography equipment including a reference mirror, thesuccessive sampling of the interference pattern being performed bysuccessively detecting light reflected or backscattered from the objectduring a movement of the reference mirror; and obtaining a plurality ofinitial image values by combining at least two of the plurality oforiginal initial image values, which were obtained at the differentdepths within the depth range of the object, to one initial image value,wherein the plurality of initial image values are located on athree-dimensional regular grid and adjacent initial image values areequally spaced apart in all three spatial dimensions, and the pluralityof initial image values define at least two two-dimensional images ofthe object in planes of the object that are spaced apart from oneanother; and interpolating the plurality of initial image values of theat least two two-dimensional initial images in three-dimensional spaceby at least one of trilinear interpolation, triquadratic interpolation,tricubic interpolation, and multi-variate interpolation in thethree-dimensional regular grid on which the plurality of initial imagevalues of the at least two two-dimensional initial images are located toobtain interpolation values that form a two-dimensional final image. 2.The method according to claim 1, wherein the plurality of initial imagevalues of the at least two two-dimensional initial images areinterpolated by at least one of the trilinear interpolation and thetricubic interpolation.
 3. The method according to claim 1, wherein theat least two two-dimensional initial images are real time images thatare acquired at a rate of at least one image per second.
 4. The methodaccording to claim 1, wherein, in a first operating mode, acquiring theat least two two-dimensional initial images in planes of the object thatare spaced apart from one another and light reflected or backscatteredby the object is detected only by a partial surface of a spatiallyresolving detector of the optical coherence tomography equipment; andchanging an optical distance of a reflector from a beam splitter of theoptical coherence tomography equipment by an optical path that is largerthan a mean wavelength λ₀ of light injected into the optical coherencetomography equipment.
 5. The method according to claim 1, wherein, in asecond operating mode, acquiring the at least two two-dimensionalinitial images of the object in two planes of the object that are spacedapart from one another during changing of an optical distance of areflector from a beam splitter of the optical coherence tomographyequipment; and detecting the light reflected from the object a pluralityof times by detector elements of a detector; wherein the changing of theoptical distance of the reflector from the beam splitter is at mostforty times a mean wavelength λ₀ of light injected into the opticalcoherence tomography equipment.
 6. The method according to claim 5,wherein the two planes extend at different depths in the object aboveand below a mean depth in the object.
 7. The method according to claim6, wherein the two planes have a same distance from the mean depth. 8.The method according to claim 6, wherein the mean depth or thedifferent, respectively, of the planes extending above or below the meandepth in the object are set by the optical distance of the reflectorfrom the beam splitter.
 9. The method according to claim 8, wherein thechanging of the optical distance of the reflector from the beam splitterof the optical coherence tomography equipment is changed by an opticalpath that is larger than the mean wavelength λ₀ of the light injectedinto the optical coherence tomography equipment.
 10. The methodaccording to claim 6, wherein a spacing of the two planes relative toone another corresponds to a same spacing of the plurality of initialimage values in all three spatial dimensions.
 11. The method accordingto claim 1, wherein spacing of two planes of the at least twotwo-dimensional images planes relative to one another corresponds to asame spacing of the plurality of initial image values in all threespatial dimensions.
 12. The method according to claim 1, wherein thedepth range of the object in the at least one dimension corresponds to aresolution of the optical coherence tomography equipment.
 13. The methodaccording to claim 1, wherein the at least two two-dimensional initialimages of the object are acquired in planes of the object that extendparallel to one another.
 14. The method according to claim 1, whereinthe combining step reduces a number of the plurality of initial imagevalues in a direction of the depth range of the object.
 15. The methodaccording to claim 1, wherein the depth range of the object in the atleast one dimension corresponds to an axial resolution or a depthresolution of the optical coherence tomography equipment.
 16. A systemfor performing optical coherence tomography, the system comprising: anoptical coherence tomography equipment including an interferometerincluding a reference mirror, a light source that generates light andinjects the light into the interferometer to illuminate an object, and adetector that detects light that is reflected or backscattered from theobject; and a processor that: obtains a plurality of initial imagevalues using the optical coherence tomography equipment by combining atleast two of a plurality of original initial image values, which wereobtained by successive sampling of an interference pattern at differentdepths within a depth range of the object, into one initial image value,the successive sampling of the interference pattern being performed bysuccessively detecting the light reflected or backscattered from theobject during a movement of the reference mirror, wherein the pluralityof initial image values are located on a three-dimensional regular gridand adjacent initial image values are equally spaced apart in all threespatial dimensions, and the plurality of initial image values define atleast two two-dimensional images of the object in planes of the objectthat are spaced apart from one another; and that interpolates theplurality of initial image values of the at least two two-dimensionalinitial images in three-dimensional space by at least one of trilinearinterpolation, triquadratic interpolation, tricubic interpolation, andmulti-variate interpolation in the three-dimensional regular grid onwhich the plurality of initial image values of the at least twotwo-dimensional initial images are located to obtain interpolationvalues that form a two-dimensional final image.
 17. The system accordingto claim 16, wherein the at least two two-dimensional initial images ofthe object are acquired in planes of the object that extend parallel toone another.